The present invention relates generally to the field of detector arrays and more specifically to the field of X-ray detector arrays in CT scanner applications.
Computed tomography (CT) X-ray scanners are used in a variety of applications. For example, such scanners are used in X-rays in medical diagnostic applications and for X-ray baggage inspection in airport security systems. For the most part, a CT scanner includes at least one X-ray source and a series of X-ray detectors. The detectors are disposed diametrically opposite the X-ray source on a rotating disk. During rotation the X-ray source emits X-rays which pass through the object being scanned and ultimately impinge on the detectors. Given that the original signal characteristics of the emitted X-rays are known, by measuring the attenuated signals arriving at the detectors, the electronics determines the density distribution in the object. Algorithms for determining an object's density based on such signal measurements are well known in the art.
In most CT systems, the X-ray detectors each first translate the received X-ray signal into an optical signal and then translate the optical signal into an electrical signal, which is processed by electronics forming part of each system. The electronics then process the electrical signals in accordance with specific application algorithms. A detector of this type often is made of a light emitting scintillating element (e.g., a scintillator crystal) paired with an optical detector or "photo-detector" (e.g., a photodiode). The scintillator crystal receives the X-ray signal and responsively generates an optical signal (e.g., blue light). The optical signal from each crystal is then detected by its corresponding photodiode, which responsively generates an electrical signal that is a function of the original X-ray flux received by the scintillator crystal. A typical detector array takes the form of a two-dimensional (i.e., m.times.n) array of detectors, or m.times.n scintillator crystal and photodiode pairs. It is important in such a detector array that light emitted from one scintillator crystal is not sensed by adjacent photodiodes which are adjacent to the intended photodiode with which the light emitting scintillator crystal is paired. Such light leakage, referred to as "optical cross-talk", causes inaccuracies in the measurement (e.g., noisy signals, erroneous detections by adjacent detectors, artifacts, etc.) and, therefore, in the X-ray system overall.
One X-ray detector array of the prior art is described in pending U.S. patent application Ser. No. 08/948450, assigned to Analogic Corporation of Peabody, Mass. and incorporated herein by reference. As shown in FIGS. 1-4 of the present application, the prior art X-ray detector system includes a large number of relatively small individual detector elements, or scintillator crystal/photodiode pairs, arranged in a two-dimensional (2-D) array. The detector array incorporates a multi-functional structure comprising a set of alignment grids which function both to align each individual scintillator crystal with a corresponding photodiode and also to isolate the individual photodiode/crystal pairs from one another to prevent optical cross-talk. Overall, the detector array is substantially stable under the typical operating conditions of the CT scanning system, which include vibration and/or temperature fluctuations.
As illustrated schematically in FIG. 1, a substrate 12 provides the basic structural support of the prior art detector array. Photodiodes 14 are arranged on the substrate in a 2-D array. As an example, a single m.times.n array may comprise 72 photodiodes arranged in six rows of twelve photodiodes each (i.e., a 6.times.12 array). The substrate 12 also includes a signal transmission arrangement 16 for transmitting electrical signals generated by the photodiodes to a signal processing subsystem 20 for image reconstruction. The signal transmission arrangement 16 can include electrically conductive circuit paths printed into the substrate, or an electrically conductive interconnect layer 17 attached to the substrate. Electrically conductive leads 19 from each photodiode to one or more of the paths complete an electrical connection between each photodiode and the signal processing means 20.
A scintillator crystal assembly 18 is positioned over the photodiode array and includes a number of scintillator crystals 22 and alignment grids for arranging the crystals in a 2-D array which corresponds to the photodiode array. Each of the scintillator crystals 22 is substantially aligned and interfaced with a corresponding photodiode 14 and is also substantially optically isolated from surrounding crystals. As shown in FIG. 2, at least one of the alignment grids 24 is substantially planar and includes a number of cells or openings 26. Each of the cells 26 is of a sufficient dimension to receive and substantially align with a scintillator crystal 22, as shown in FIG. 3. Another alignment grid 28 is optically opaque and substantially rigid 2-D grid, having a significant thickness relative to grid 24. It also includes a number of cells 26' corresponding to the cells of the first alignment grid 24. Each of the cells of the alignment grid 28 is substantially aligned with a corresponding cell of the planar alignment grid 24 and thus with a scintillator crystal 22. Optical opacity and dimensional stability are critical features of the alignment grids.
The alignment grids 24 and 28 provide a structural framework for the scintillator crystals 22 in the detector array which ensures the correct alignment of the crystals with corresponding photodiodes 14 and provides dimensional stability to the crystal assembly. As shown in FIGS. 2 and 3, the cells 26, 26' of the respective alignment grids 24, 28 are each sized to accommodate and align a single scintillator crystal with a corresponding photodiode 14. The 2-D alignment grid 28 includes walls which extend above the photodiode array and electrical interconnections 17 on the substrate. The walls are positioned directly in between photodiode detectors and establish individual wells or cells for each scintillator crystal.
The cell width of the 2-D grid is sufficiently large to accommodate the bonding of wire 19 from the detection side of photodiode 14, the traversal of the wire 19 down the side of the photodiode, and bonding of the wire to the electrical layer 17. Because the wire leads 19 from the photodiodes may be relatively fragile, effort is taken to protect them from damage. The 2-D alignment grid 28 additionally serves as a standoff between the photodiode array and a support for the scintillator crystal assembly so that the crystals 22 cannot rest directly on corresponding underlying photodiodes 14 and wire leads 19, which would likely cause damage to the relatively fragile wire leads 19. Therefore, the height and width of grid 28 is at least as great as the height and width of the combined photodiode 14 and wire lead 19, as shown in FIG. 4.
As is also shown in FIG. 4, the scintillator crystals 22 are surrounded on all sides, other than the side closest to a corresponding photodiode 14, by an optically reflective material 30, like paint, foil, or surface deposition layers. The region between a scintillator crystal 22 and a corresponding photodiode 14 is filled with an optically transmissible medium 34 (e.g., epoxy) to facilitate transmission of light from the crystal to the photodiode.
As will be appreciated by those skilled in the art, manufacture of the prior art X-ray detector array tends to be complex and labor intensive, due to the precautions necessary to insure its reliable construction. For example, each wire 19 must be bonded to the photodiode, and carefully looped, such that the wire when bonded to a circuit path on the substrate 12 does not touch the photodiode wall or grid 28. Additionally, alignment grid 28 is itself fragile and vulnerable to breakage, given that grid 28 tends to be made of a brittle material such as glass or ceramic. Also, the loop of wire 19 from the top of photodiode 14 makes the wire vulnerable to damage or displacement and requires that the scintillator crystal array be sufficiently raised above the photodiode array to provide clearance for the wire. Such separation increases the likelihood of optical cross-talk and, thus, might impact the accuracy of the X-ray system overall. Additionally, to compensate for the traversal of wire 19 along the side of the photodiode as it makes its way to the substrate 12, the grid 28 must be precisely fitted to the photodiode array. The errors in fitting might compromise the physical integrity of the photodiode or wire bonding.
Therefore, it is an object of the present invention to provide a 2-D photo-detector array with reduced optical cross-talk, improved structural integrity, and a simplified structure leading to lower manufacturing costs. It is a further object of the invention to provide an improved X-ray detector having greater accuracy, durability, and reliability, which is substantially less costly to manufacture than prior art X-ray detectors. And, it is yet another object of the present invention to provide an improved CT X-ray scanner system which achieves these same benefits.